1. Field of the Invention
The present invention pertains to a system and method for determining a respiratory condition of a patient and to a system and method for controlling a ventilation system based on the determined condition, and, in particular, to a system and method that non-invasively measures a patient's elastance and resistance and to a ventilator that employs such a method to measure elastance and resistance during ventilation so that the ventilatory assistance provided to the patient by the ventilator is automatically adjusted to suit the needs of the patient.
2. Description of Related Art
The related art will be described with reference to the following patents and other publications, the disclosures of which are hereby incorporated by reference in their entireties into the present disclosure. Throughout the description of the related art, these references will be cited by the first-named author and the year of publication, e.g., Jackson, 1974.
M. Franetzki et al, U.S. Pat. No. 4,051,843 (1977); U.S. Pat. No. 4,022,193 (1977); and U.S. Pat. No. 4,122,839 (1978).
M. M. Grunstein, U.S. Pat. No. 4,802,492 (1989).
P. H. Vooren, U.S. Pat. No. 4,259,967 (1981).
Y. Yoshitsugu, European Published Patent Application 0 521 515 A1 (1993).
P. J. Chowiency, C. P. Lawsom, et al, U.S. Pat. No. 5,233,998 (1993).
Bates, J. H. T., Daroczy, B., and Hantos, Z., "A Comparison of Interrupter and Forced Oscillation Measurements of Respiratory Resistance in the Dog," Journal of Applied Physiology, Vol. 72, Iss. 1., pp. 46-52 (1992).
Bates, J. H. T., Decramer, M., Zin, W. A., Harf, A., "Respiratory Resistance with Histamine Challenge by Single-breath and Forced Oscillation Methods," Journal of Applied Physiology, Vol. 61, No. 3, pp. 873-80 (1986).
Bates, J. H. T., Baconnier, P., Milic-Emili, J., "A Theoretical Analysis of Interrupter Technique for Measuring Respiratory Mechanics," Journal of Applied Physiology, Vol. 64, No. 5, pp. 2204-14 (1988).
Calhoun, Karen H., House, William, et al, "Normal Nasal Airway Resistance in Noses of Different Kinds and Shapes," Otolaryngology Head and Neck Surgery, Vol. 103, No. 4, pp. 605-9 (1990).
Chatburn, Robert L., "A New System for Understanding Mechanical Ventilators," Respiratory Care, Vol. 36, No. 10, pp. 1123-55 (1991).
Chiang, S. T., Green, J., Gao Y. C., "Determination of Total Respiratory Resistance in Health and Disease by Added External Resistance," Chest, Vol. 93, pp. 537-40 (1988).
Chowienczyk, P. J., Lawson, C. P., et al, "A Flow Interruption Device for Measurement of Airway Resistance," European Respiratory Journal, Vol. 4, pp. 623-628, (1991).
Daroczy, B., Hantos, Z., "Generation of Optimum Pseudorandom Signals for Respiratory Impedance Measurements," International Journal of Biomedical Computation, Vol. 25, pp. 21-31 (1990).
Frank, N. R., Mead, J., Whittenberger, "Comparative Sensitivity of Four Methods for Measuring Changes in Respiratory Flow Resistance in Man," Journal of Applied Physiology, Vol. 31, No. 6 (December, 1971).
Green, J., Chiang, S. T., Yang Y. C., "Improved Computation of Respiratory Resistance as Measured by Transiently Increased Resistance," Medical & Biological Engineering & Computing, Vol.28, pp. 50-53 (1990).
Gimeno, F., van der Weele, L. Th., "Variability of Forced Oscillation (Siemens Siregnost FD5) Measurements of Total Respiratory Resistance in Patients and Health Subjects," Annals of Allergy, Vol. 71, pp. 56-60 (July, 1993).
Hantos, Z., Daroczy, B., Suki, B., "Forced Oscillatory Impedance of the Respiratory System at Low Frequencies," Journal of Applied Physiology, Vol. 60, pp. 123-32 (1986).
Jackson, A. C., Milhom, H. T., and Norman, J. R., "A Reevaluation of the Interrupter Technique for Airway Resistance Measurement," Journal of Applied Physiology, Vol. 36, No. 2 (December, 1974).
Lutchen, Kenneth, Yang, Kun, Kaczka, David W., "Optimal Ventilation Waveforms for Estimating Low-Frequency Respiratory Impedance," Journal of Applied Physiology, Vol. 75, Iss. 1, pp. 478-88 (1993).
Lutchen, Kenneth., Kaczka, David W., Suki, Bela, "Low-frequency Respiratory Mechanics Using Ventilator-driven Forced Oscillations," Journal of Applied Physiology, Vol. 75, No. 6, pp. 2549-60 (1993).
Mayewski, Raymond J., Hyde, Richard W., "Measurement of Static Pressure-Volume Relationships of the Lung and Thorax," The Selective and Comprehensive Testing of Adult Pulmonary Function, E. Leslie Chusid, ed., Futura Publishing Co., New York (1983).
Morozoff, Paul E., Evans, Ron W., "Real-Time Display of Flow-Pressure-Volume Loops," Biomedical Instrumentation & Technology (July/August, 1992).
Neild, J. E., "The Repeatability and Validity of Respiratory Resistance Measured by the Forced Oscillation Technique," Respiratory Medicine, Vol. 83, pp. 111-18 (1989).
Petak, F., Hantos, Z., Adamicza, A., "Partitioning of Pulmonary Impedance: Modeling vs. Alveolar Capsule Approach," Journal of Applied Physiology, Vol. 75, No. 2, pp. 513-521 (1993).
Romero, P. V., Sato, J., Shardonfsky, F., "High-frequency Characteristics of Respiratory Mechanics Determined by Flow Interruption," Journal of Applied Physiology, Vol. 69, No. 5, pp. 1682-88 (1990).
Suki, Bela, Lutchen, Kenneth R., "Pseudorandom Signals to Estimate Apparent Transfer and Coherence Functions of Nonlinear Systems: Applications to Respiratory Mechanics," IEEE Transactions on Biomedical Engineering, Vol. 39, No. 11 (November, 1992).
Suki, B., Hantos, Z., "Nonlinearity and Harmonic Distortion of Dog Lungs Measured by Low-Frequency Forced Oscillations," Journal of Applied Physiology, Vol. 71, pp. 69-75 (1991).
Suki, B., Peslin, R., Duvivier, C., "Lung Impedance in Health Humans Measured by Forced Oscillations from 0.01 to 0.1 Hz," Journal of Applied Physiology, Vol. 67, No. 4, pp. 1623-29 (1989).
To understand how a machine can be controlled to replace or supplement the natural function of breathing, it is necessary to understand the mechanical nature of the respiratory system. The study of the mechanical behavior of the respiratory system requires analyzing the elastance and resistance properties of the patient's pulmonary system, which includes the airways, lung and thoracic cage. In clinical practice, respiratory resistance R.sub.rs and elastance E.sub.rs are essential information necessary to describe the behavior of the lung and the chest wall in health and disease states, and, in particular, to describe characteristics of that behavior, such as inspiratory vital capacity (IVC) and the forced expiratory volume in one second (FEV1). Furthermore, the use of state-of-the-art mechanical ventilation techniques, such as proportional assist ventilation (PAV), which is disclosed in U.S. Pat. Nos. 5,107,830 and 5,044,362 both to Younes, the contents of which are also incorporated herein by reference, requires knowledge of the patient's respiratory resistance and elastance.
Measuring the respiratory resistance and elastance of a spontaneously breathing patient is not a simple task. Conventional techniques for measuring resistance and elastance are somewhat invasive in that they are performed in a clinical or hospital setting and require placing a device for measuring esophageal pressures, such as an esophageal balloon, within the patient. Therefore, R.sub.rs and E.sub.rs are typically not measured on a routine basis. In order to perform these measurements more routinely, there is a need for an efficient and reliable technique that is as non-invasive as possible and requires little or no patient cooperation for spontaneously obtaining R.sub.rs and E.sub.rs, especially inspiratory R.sub.rs and E.sub.rs.
Respiratory mechanics takes into consideration the forces, displacement, rate of change (first time derivative) of displacement, and acceleration (second time derivative) of displacement. In respiratory physiology, force is measured in terms of pressure P, displacement is measured as volume V, rate of change of displacement is measured as flow V (first time derivative), and acceleration of displacement is measured as the rate of change of flow V (second time derivative of displacement). Particularly relevant to assisted breathing is the pressure P necessary to cause a flow of gas V, thereby increasing the volume of the lungs V against the inertial force of the respiratory system caused by the rage of change of flow V.
Over the course of a breathing cycle, i.e., one inspiration and one expiration, pressure P(t) (typically measured in cm H.sub.2 O), volume V(t) (typically measured in liters), flow V (t) (typically measured in liters/second) and rate of change of flow V (t)(typically measured in liters/second.sup.2) all change with time. The total force, i.e., pressure, necessary to expand the lungs and chest wall must overcome the following three different forces: inertial force, resistive force, and elastic recoil force, all of which are developed by the respiratory system and oppose its expansion. A mathematical model, i.e., the equation of motion, for the respiratory system describes the relation among the pressure, flow and volume as follows: EQU P.sub.aw (t)+P.sub.mus (t)=IV(t)+R.sub.rs V(t)+E.sub.rs V(t) (1)
In this equation, P.sub.aw (t) is the ventilator pressure applied at the airway opening. Muscle pressure P.sub.mus (t) is the imaginary transrespiratory pressure (airway pressure--body surface pressure) generated by the ventilatory muscles to expand the thoracic cage and lungs. Muscle pressure P.sub.mus (t) is not directly measurable.
Elastic force E.sub.rs V(t) is the force with which the respiratory system attempts to recoil after deflation. The elastic force is generated by the lung and thorax elastic supporting structures. E.sub.rs is defined as the change in distending pressure per change in volume and is the reciprocal of compliance and is expressed in units of cm H.sub.2 O/liter. The total static respiratory recoil volume pressure P.sub.rs is given by the sum of pressure P.sub.l developed across the lungs and pressure P.sub.cw developed across the chest wall: EQU P.sub.rs =P.sub.l +P.sub.cw. (2)
Because volume change V.sub.rs in the respiratory system is given by the sum of volume change V.sub.l in the lungs and volume change V.sub.cw in the chest wall, the total respiratory elastance E.sub.rs is given by the sum of lung elastance E.sub.l and chest wall elastance E.sub.cw : EQU E.sub.rs =E.sub.l +E.sub.cw. (3)
Total respiratory elastance E.sub.rs is dependent on factors such as lung size, the sex of the patient, the growth and aging of the patient, the resting positions of the lungs in the thorax, and gravitational (positional) effects. The dynamic elastance has a marked deviation from the static elastance because of uneven time constants in the airways and lung parenchyma (See Mayewski, 1983).
Resistive force R.sub.rs V(t) is the force exerted by the movement of gas and tissue elements in the lungs and thorax that oppose movement of the lungs and thorax. Total respiratory resistance R.sub.rs is determined by dividing the pressure gradient between the airway opening and the body surface of the chest cage required to overcome non-elastic and non-inertial factors by flow. The pressure gradient used in the measurement of R.sub.rs includes the sum of the pressure necessary to move air through the airways (which gives R.sub.aw), the pressure necessary to change the shape of the lung tissues (which gives tissue viscous resistance R.sub.visc) and the pressure necessary to move the chest wall and the diaphragm (which gives R.sub.wall). R.sub.rs is expressed in units of cm H.sub.2 O/(liter/second), or cm H.sub.2 O.multidot.sec/liter, and is given by the following formula: EQU R.sub.rs =R.sub.aw +R.sub.visc +R.sub.wall (4)
Inertial force IV(t) is the force introduced by the inertial property of the respiratory system. It is proportional to the rate of change of flow. Under normal circumstances, this force is usually negligible. However, the effect of the inertial force increases with increases in the patient's ventilation rate.
Equation (1) provides a dynamic model in which pressure, flow and volume are all measured relative to their baseline values (i.e., their values at the end of expiration). The pressure that causes inspiration is measured as the change in airway pressure above positive end-expiratory pressure (PEEP). The volume is measured as the change in lung to volume above the functional residual capacity (FRC). Flow is measured with respect to its end-expiratory value, which is usually zero.
The parameters in Equation (1) are not necessarily constant. In fact, the mechanical behavior of the respiratory system has been characterized as nonlinear. Almost every mechanical aspect of lung behavior can exhibit nonlinear characteristics. The pressure-area behavior of the airway walls, the pressure-volume behavior of the lung parenchyma and the pressure-flow behavior of the airway gas are all well documented as being nonlinear. The variables known to change resistance and elastance are very complex, including flow rate, lung volume, points in the ventilatory cycle and ventilatory rate. However, the dominant factors in the nonlinear properties of R.sub.rs and E.sub.rs are flow and volume, respectively. Thus, R.sub.rs can be expressed approximately as a function of flow, R.sub.rs (V(t)). Likewise, E.sub.rs can be expressed approximately as a function of volume, E.sub.rs (V(t)). As a further simplification, the following first order equations may be used to reflect the nonlinear factors: EQU E.sub.rs =E.sub.rs0 +E.sub.rs1 V(t) (5) EQU R.sub.rs =R.sub.rs0 +R.sub.rs1 V(t) (6)
where E.sub.rs0 and R.sub.rs0 are constant terms and E.sub.rs1 and R.sub.rs1 are first-order terms. With this approximation, the equation of motion can be expressed in the first order as follows: EQU P.sub.aw (t)+P.sub.mus (t)=(E.sub.rs0 +E.sub.rs1 V(t))V(t)+(R.sub.rs0 +R.sub.rs1 V(t))V(t)+IV(t) (7)
Many other models have been developed in recent years, including a sophisticated physiological model, that reflects tissue viscoelasticity as well as the inertial effects of the airways and branching networks. However, the use of nonlinear models precludes application of many powerful concepts typically employed in a clinical investigation of respiratory mechanics, such as the use of frequency-domain analysis, Bode diagrams and multilinear regression. In most cases, it is acceptable to consider R.sub.rs and E.sub.rs as constant and to use the following two-element linear model: EQU P.sub.aw (t)+P.sub.mus (t)=E.sub.rs0 V(t)+R.sub.rs0 V(t) (8)
A study of respiratory mechanical properties is an important area of interest to respiratory care professionals. In accordance with the analysis of mechanics and breathing, respiratory mechanics can be assessed if it is possible to measure P.sub.aw (t), V(t) and P.sub.mus (t). The first two variables are easily measured by means of sensors located at the airway opening. However, there is believed to be presently no known direct method of non-invasively measuring P.sub.mus (t) under dynamic conditions.
There are other situations in which it is it important to know the patent's respiratory mechanics. For example, in order to implement proportional assist ventilation, which is a synchronized partial ventilation method that amplifies patient respiratory effort to deliver pressure to the patient in proportion to the patient's instantaneous effort, a knowledge of patient's respiratory mechanics is required. The respiratory mechanics, such as resistance and elastance, are used in a PAV system to determine the proper level of flow and volume support. For a ventilator supported patient, such parameters constantly vary because of different physical and pathological conditions. Therefore, it is important to be able to continuously or periodically determine these parameters while minimizing the obtrusiveness of the measurements required to do so.
In short, PAV requires accurate resistance and elastance values to maintain optimal flow and volume support so that the pressure support truly accommodates the patient's breathing effort. Compared with other ventilation modes, PAV requires detailed information on respiratory mechanics and more interaction with the patient. Such interaction is preferably performed on an ongoing basis, because, as noted above, respiratory mechanics are variable for most patients.
Many noninvasive respiratory mechanics measurement techniques have been developed. In general, these measurement techniques can be divided into the following five categories: interrupter/occlusion, variable external resistance, time constant, multi-linear regression and forced oscillation. However, these conventional clinical techniques for determining resistance and elastance are cumbersome and cannot be performed easily on a ventilator supported patient. Furthermore, several of these techniques require manually implemented procedures and cannot be performed using most ventilators, especially ones that exhibit system leak.
The interrupter/occlusion method estimates the mean alveolar pressure. See Jackson, 1974. This measurement method entails providing a rapid occlusion, e.g., approximately 0.1 second, in the breathing circuit during a normal breathing cycle. This technique assumes that during the occlusion, the alveolar pressure and the pressure at the airway opening equilibrate so rapidly that the net movement of the rib cage and the diaphragm does not change intrapleural pressure appreciatively, although continued respiratory effort is still present. The pressure measured at the airway opening immediately after equilibration is used to estimate the alveolar pressure just prior to the occlusion.
The interrupter/occlusion method is the most common clinical practice for estimating lung elastance. The occlusion is usually performed at the beginning of an exhalation. Upon providing the occlusion, the pressure at the airway opening increases and plateaus in about 250 ms, when the respiratory muscles are completely relaxed. The plateau pressure equilibrates with the respiratory elastic recoil force. Because the occlusion is provided at the beginning of exhalation, the total air volume in the respiratory system equals to the tidal volume V.sub.tidal plus the functional residual capacity. The elastance E.sub.rs can be determined if pressure and volume are known.
For a non-leak system, clinicians manually block the exhalation path using their hands, an exhaustion valve or a shutter. In an open circuit system, however, the system leak, e.g., exhalation or exhaust port, is typically located very close to the patient's airway. Therefore, it is not practical to insert a shutter between the leak and the airway opening. A commercially available flow interruption device is taught by U.S. Pat. No. 5,233,998 for measuring airway resistance, but not total respiratory resistance, R.sub.rs. This technique is difficult to implement on a ventilators with a system leak for the reasons noted above, namely it is not practical to provide the occlusion between the leak and the patient's airway.
The variable external resistance technique is reported to measure R.sub.rs by using a rapid and brief increase in external resistance (R.sub.ext). See M. Franetzki, 1977; Chiang, 1988; and Green, 1990. This measurement method is based on the assumption that while the external resistance is in series with R.sub.rs, the changes in muscle pressure (P.sub.mus) and elastic force are negligible. This technique, however, is incomplete in that it ignores the possible effect caused by the inertial factor of the respiratory system. It is also difficult to incorporate this technique into a ventilator because it requires measuring linear resistance, which cannot be done readily using conventional ventilators.
The time constant method is used to estimate resistance and elastance during an expiratory phase by examining how exhalation decays. See, e.g., Grunstein, U.S. Pat. No. 4,802,492, 1989. However, studies have shown that inspiratory resistance and elastance values are different from the expiratory resistance and elastance values. Therefore, this technique is not well suited to measure or estimate inspiratory resistance and elastance.
The multi-linear regression method is used to estimate resistance of an anesthetized patient whose muscle pressure is eliminated. Accordingly, this technique is not applicable for an active patient.
The forced oscillation technique applies an oscillated pressure at the patient's airway opening. It has become a very popular means for scientists to study the respiratory system. Some studies suggest that the oscillation frequency should be set around 6-7 Hz because that is the resonant frequency of the respiratory system in humans. See Frank, 1971. Studies have also reported that reliable estimates of resistance cannot be expected at frequencies lower than 2-4 Hz, especially for patients whose breathing pattern is relatively rich in harmonics, such as vigorously breathing children or an obstructed patient. See, e.g., Daroczy, 1990 and Hantos, 1986.
In clinical applications, due to the expensive and bulky instrument required, this technology was not widely used until Siemens introduced the Siregnost FD-5 portable oscillometer in the early 1980's. See Gimeno, 1993. After a clinical study, Neild et al. (1989) proved the repeatability and validity of a derived measurement of R.sub.rs obtained with the Siregnost FD-5. One of the reasons the forced oscillation technique has become a widely used method for measuring the total respiratory resistance is the fact that the patient's cooperation can be kept at minimum. Most forced oscillation devices use loudspeakers in enclosures or linear motor pumps as a high frequency pressure oscillation source to produce controlled perturbations in the airway. In addition, the existing forced oscillation technique use a sequenced pressure oscillation during the patient's entire inspiration phase. Therefore, the conventional forced oscillation technique is not practical for ventilator supported patients on an ongoing basis.